Positron Emission Tomography (PET) is a nuclear imaging technique used in the medical field to assist in the diagnosis of diseases. In PET, short-lived positron-emitting isotopes, referred to as radiopharmaceuticals, are injected into a patient. When these radioactive drugs are administered to a patient, they distribute within the body according to the physiologic pathways associated with their stable counterparts. For example, the radiopharmaceutical .sup.18 F-labeled glucose, known as fluorodeoxyglucose or "FDG", can be used to determine where normal glucose would be used in the brain. Other radioactive compounds suitable for PET scanning include .sup.11 C-labeled glucose, .sup.13 N-labeled ammonia and .sup.15 O-labeled water.
As the FDG or other radiopharmaceutical isotopes decay in the body, they discharge positively charged particles called positrons. Upon discharge, the positrons encounter electrons, and both are annihilated. As a result of each annihilation event, gamma rays are generated in the form of a pair of photons approximately 180 degrees (angular) apart. These occurrences can be mapped within the patient's body, thus allowing for the quantitative measurement of metabolic, biochemical and functional activity in living tissue. More specifically, PET images (often in conjunction with an assumed physiologic model) can be used to evaluate a variety of physiologic parameters such as glucose metabolic rate, cerebral blood flow, tissue viability, oxygen metabolism and in vivo brain neuron activity.
Positron Emission Tomography (PET) has gained significant popularity in nuclear medicine because of the ability to non-invasively study physiological processes within the body. PET allows the physician to examine the whole patient at once by producing pictures of many functions of the human body unobtainable by other imaging techniques. In this regard, PET displays images of how the body works (physiology or function) instead of simply how it looks. PET is considered the most sensitive, and exhibits the greatest quantification accuracy, of any nuclear medicine imaging instrument available at the present time. Applications requiring this sensitivity and accuracy include those in the fields of oncology, cardiology and neurology.
As noted, PET data acquisition occurs by detection of both photons emitted from the annihilation of the positron in a coincidence scheme. Due to the approximate 180 degree angle of departure from the annihilation site, the location of the two detectors registering the "event" define a chord passing through the location of the annihilation.
By histogramming these lines of response (the chords), a "sinogram" is produced that may be used by a process of back-projection to produce a two dimensional image of the activity. Detection of these lines of activity is performed by a coincidence detection scheme. A valid event line is registered if both photons of an annihilation are detected within a coincidence window of time. Coincidence detection methods ensure (disregarding other second-order effects) that an event line is histogrammed only if both photons originate from the same positron annihilation.
Another tomography diagnostic system has been developed, known as single photon emission computed tomography (SPECT). The SPECT apparatus is similar to a PET system, except that instead of analyzing photon pairs, the SPECT system analyzes single photons emitted from the positron annihilations detected within the patient. It is now possible for a single tomography system to employ both PET and SPECT technologies.
Traditionally, positron emission tomography systems have employed discrete scintillators arranged in rings. The scintillators are comprised of materials suitable for interacting with gamma rays, including bismuth germanium crystals (BGO), lutetium oxyorthosilicate crystals (LSO), and sodium iodide crystals (NaI). Typically there are hundreds of detectors per ring, with one to 100 rings in the detector structure. Data is collected through two-dimensional acquisition of photon emission events. Recent advances in the art have permitted three-dimensional, or 3-D acquisition of data using an assortment of discrete scintillators.
The details of carrying out a PET study are given in numerous publications. Typically, the following references provide a background for PET. These are incorporated herein by reference for any of their teachings.
1. M. E. Phelps, et al.: "Positron Emission Tomography and Audiography", Raven Press, 1986; PA0 2. R. D. Evans: "The Atomic Nucleus", Kreiger, 1955; PA0 3. J. C. Moyers: "A High Performance Detector Electronics System for Positron Emission Tomography", Masters Thesis, University of Tennessee, Knoxville, Tenn, 1990; PA0 4. U.S. Pat. No. 4,743,764 issued to M. E. Casey, et al, on May 10, 1988; PA0 5. S. R. Cherry, et al.: "3-D PET Using a Conventional Multislice Tomograph Without Septa", JI. C.A.T., 15(4) 655-668; and PA0 6. C. L. Morris, et al.: "A Digital Technique for Neutron-Gamma Pulse Shape Discriminator", Nuclear Inst. and Methods, 137 (1976) 397-98.
It is desirable for the customer and, derivatively, the patient for the cost of PET scanner systems to be reduced. Cost reductions have been achieved through the use of block detectors. The block detectors allow for a reduction in the number of photomultiplier tubes (PMT) required to properly detect the gamma ray interaction location. In this arrangement, more crystals are enabled to share each PMT.
Another advance in PET is the ability to determine the depth of interaction of the gamma ray in the detector. The depth information is used to correct the line of response determination due to the penetration depth of oblique angle gamma rays on large area detectors. This parallax problem arises at the edge of the field of view on large area detectors and in ring tomographs. To determine the depth of interaction of a gamma ray in a detector, a special detector composed of two or more different decay time scintillating crystals is used. In this arrangement, the crystals are placed on top of each other from front to back. The depth of interaction of each gamma ray is determined by analyzing the light decay properties of the detected scintillations to determine the crystal in which the gamma ray interacted; i.e., at the front or the back of the crystal arrangement. The gamma ray interaction depth is used in the image reconstruction to correct the line of response parallax.
At present, there are multi-crystal tomography imaging systems which employ two scintillating crystals in the detector. These dual crystal detector arrangements are referred to as phoswitch detectors. Reasonable costs, high light output, and fast decay time scintillating crystals have made phoswitch detectors commercially and technologically practical. However, a data collection problem arises in connection with the use of dual crystal detectors. That problem is based on the different decay times incident in the two crystals. Those skilled in the art will understand that crystals having different decay times will necessarily have different optimal integration times for carrying out the signal processing process. More specifically, the optimal integration time is longer for the slow crystal than for the fast crystal. However, a longer integration time creates a circumstance wherein multiple gamma rays may be incident on the fast crystal during the integration process, causing a "pile-up" of gamma ray events. This, in turn, can cause the sampling of multiple events within the fast crystal to be misinterpreted as a single gamma ray event within the slow crystal. The result is that the image reconstruction used to correct the line of response parallax will report an incorrect position.
The problem of pile-up is compounded when phoswitch detectors are used in a block configuration. Those skilled in the art understand that annihilation events in connection with PET occur randomly. Due to the large area of the detector, multiple gamma rays may be incident on the individual detectors during the integration time, that is, the period in which gamma rays are analyzed. The use of large area detectors increases the probability that one gamma ray will interact with a scintillating crystal while the signal resulting from an earlier gamma ray is still being processed, i.e., during the "dead time."
Accordingly, it is an object of the present invention to provide a system and method for reducing pile-up errors in dual crystal tomography applications.
It is a further object of the present invention to provide a system and method for reducing pile-up errors in multi-crystal structures for PET scanning systems having block crystal configurations.
Another object of the present invention is to provide a system and method for reducing pile-up of signal data in connection with single photon emission computed tomography (SPECT) systems.